Field of the Invention
The invention concerns a method and a control device to operate a magnetic resonance system to emit or activate a pulse sequence. The invention also concerns a magnetic resonance tomography system (also shortened to magnetic resonance system or “MRT”).
Description of the Prior Art
Magnetic resonance tomography is by now a widespread technique for the acquisition of images of the inside of a living examination subject.
Atomic nuclei, of hydrogen atoms, for example, have a spin. The spin is a quantum mechanical property of atomic particles. The spin has the effect that the atomic particles are magnetic, meaning that an atomic nucleus with spin is a dipole. These spins initially act in any direction, and can be considered as a vector. Atoms with spin are present in a body to be examined, for example a human body.
In a magnetic resonance tomography system, the body to be examined is typically exposed to a relatively high basic magnetic field B0, for example of 1, 5, 3 or 7 Tesla, with the use of a basic field magnet system. The static magnetic field B0 imposes a preferential direction of the spins parallel and antiparallel to the field lines. An excess always forms in one direction, which leads to a macroscopic magnetization of the spin ensemble.
A radio-frequency field B1 is superimposed on the static magnetic field B0. This radio-frequency field normally generated by radio-frequency excitation pulses, deflects the spins out of the steady state generated by the B0 field when the radio-frequency excitation signals are in resonance with the precession frequency of the spins. The precession frequency is also called Larmor frequency and is dependent on the strength of the B0 magnetic field. The radio-frequency excitation signals excite the nuclear spins of the atoms to resonance by being flipped (deflected) by a defined flip angle relative to the magnetic field lines of the basic magnetic field.
The connection between a resonant radiated RF pulse with the field strength B1 and the flip angle α that is achieved with this is provided by the equation
                    α        =                              ∫                          t              =              0                        τ                    ⁢                                    γ              ·                                                B                  1                                ⁡                                  (                  t                  )                                            ·                                                          ⁢              d                        ⁢                                                  ⁢            t                                              (        1        )            wherein γ is the gyromagnetic ratio (which is fixed material constant for most nuclear spin examinations) and τ is the duration of action of the radio-frequency pulse.
The emission of the radio-frequency signals for nuclear spin magnetization most often takes place from a “whole-body coil” or “body coil”. A typical design of a whole-body coil is a cage antenna (birdcage antenna), composed of multiple transmission rods that—proceeding parallel to the longitudinal axis—are arranged around a patient space of the MRT in which a patient is located in the examination. The antenna rods are respectively connected capacitively with one another in a ring on their front sides. However, by now body-proximal local coils are also increasingly used for the emission of MRT excitation signals. The reception of the magnetic resonance signals normally takes place with the local coils, but in some cases alternatively or additionally with the body coil. Magnetic resonance images of the examination subject are created on the basis of the received magnetic resonance signals. Each image point in the magnetic resonance image is associated with a small body volume (what is known as a “voxel”), and each brightness value or intensity value of the image points is linked with the signal amplitude of the magnetic resonance signal that is received from this voxel.
The signals emitted upon the precession and received by the reception antennas must be spatially associable in order to enable an imaging. For this purpose, in the acquisition of the signals a spatial coding is implemented by coding gradients.
For this purpose, a gradient along a gradient direction (x, y, z) is applied by gradient coils. The magnetic field B0 thereby increases linearly in the direction of the superimposed gradient. The precession of the nuclear spins is accordingly different along the gradient direction, spinning slower at some points and faster at others. They therefore show resonance at different frequencies. A spatially selective excitation of the nuclear spins is possible by the superimposed gradient field.
The control of the magnetic resonance system for acquisition of the raw data takes place according to a “pulse sequence”. A “pulse sequence” (also shortened to just “sequence”) is a combination of radio-frequency pulses B1 and magnetic gradient fields Bgrad of defined frequency or strength, which are switched on and off multiple times in a predetermined order every second.
In many sequences, a radio-frequency pulse B1 of suitable frequency (Larmor frequency), known as the 90° excitation pulse, is present at the beginning. By this pulse, the magnetization of the spins is deflected by 90° transverse to the external magnetic field B0, so these nuclei begin to spin around the original axis. As in the case for a top that is jolted, this movement is called precession.
The radio-frequency signal that thereby arises can be measured outside of the body. It decreases exponentially because the proton spins get out of “time” (“dephase”) and increasingly destructively superimpose. This depends on the chemical environment of the hydrogen and is different for every type of tissue.
By radiating a suitable 180° rephasing radio-frequency pulse, the effect can be produced that a portion of the dephasing is canceled again at the point in time of the measurement, so that more spins are again in the same phase.
In order to be able to associate individual volume elements (voxels) with the signals, a spatial coding is generated with linearly spatially dependent magnetic fields (gradient fields). Use is made of the fact that the Larmor frequency for a specific particle is dependent on the magnetic flux density (the stronger the field portion orthogonal to the direction of the particle's angular momentum, the higher the Larmor frequency).
This can proceed as follows for a typical 2D magnetic resonance pulse sequence:
A first gradient (in the z-direction, for example) is present at the excitation and ensures that only a single slice of the body possesses the matching Larmor frequency, thus that only the spins of this slice are deflected (slice selection gradient).
A second gradient (in the x-direction, for example) is switched at right angles to the two others during the measurement and ensures that the spins of each image column have a different precession velocity, thus emit a different Larmor frequency (readout gradient, frequency coding gradient).
A third gradient (in the y-direction, for example) transversal to the two others is activated briefly after the excitation and produces a controlled dephasing of the spins such that the precession of the spins has a different phase position in each image line (phase coding gradient).
All three gradients z, x, y together thus produce here a coding of the signal in three spatial planes. The received signal belongs to a defined slice of the body and includes a combination of frequency coding and phase coding.
As described above, in 2D magnetic resonance pulse sequences a spatial coding takes place in two directions or dimensions. Therefore, respective image information or raw data is read out for a very thin slice. The slice is selected in advance.
In 3D magnetic resonance pulse sequences, the spatial coding takes place in three directions or dimensions. Therefore, image information or raw data are read out for an entire volume, known as a “slab”. During the subsequent acquisition of the raw data from the excited slab, usually the readout takes place with frequency coding in one direction and a phase coding taking place in the other two directions.
The raw data are entered into a memory at data points organized as a matrix, known as k-space. The values in k-space are mathematically in frequency domain that is related to a partial domain that includes the subject magnetization, by a Fourier-conjugate, which means that the data in the frequency domain are converted into the data in the spatial domain by a Fourier transformation. The axes of k-space designate spatial frequencies. k-space has a unit that is inverse to the distance, for example 1/cm. In 3D tomography, k-space is also three-dimensional. A two-dimensional image or a three-dimensional image volume is then reconstructed with a (two- or three-dimensional) Fourier transformation from the raw data in two-dimensional or three-dimensional k-space.
Static magnetic field differences contribute to an expansion of the spins upon relaxation. With spin echo sequences, this expansion is canceled by a refocusing pulse or, respectively, by a series of refocusing pulses. If multiple refocusing pulses (normally 180° pulses) follow in succession, multiple spin echoes arise, generated by one multiecho sequence. The writing into k-space depends on, among other things, the desired contrast. The earlier echoes—meaning the echoes with a smaller position number—are often initially written into central k-space.
SPACE (Sampling Perfection with Application optimized Contrast using different flip angle Evolutions) is one example of a three-dimensional turbo spin echo sequence method (more precisely a single slab 3D turbo spin echo method) that can have very long echo trains. For example, a long echo train includes between forty and multiple hundreds of echoes; even a thousand echoes are possible, for instance. For a “provided signal development” (prescribed signal evolution), the flip angle of the refocusing pulses in an echo train is adapted to the properties (T1 and T2) of the different tissue types. A variable flip angle curve (flip angle evolution) is obtained. A desired signal strength is generated for different types of tissue. For example, a desired contrast can therefore be generated.
The exciting radio-frequency signal or the exciting radio-frequency pulse receives a defined bandwidth of neighboring frequencies around a center frequency. In this way, a desired region along the gradient direction can be excited.
In nearly all molecules, multiple hydrogen atoms are bound at different positions. Different positions mean different chemical (and therefore for the most part also different magnetic) environments. The local magnetic field gradient is hereby reduced or, respectively, increased; the resonance frequencies of the bound protons are somewhat lower or higher than the typical Larmor frequency.
The nuclear spins in the body tissue thus have no uniform precession frequency in the magnetic field, but rather differ according to their chemical environment for different tissue types. This is typically designated as a chemical shift. Fat has multiple peaks in the spectrum, but one is more strongly pronounced and delivers a high signal to the imaging. The chemical shift between the main peak of adipose tissue and water is approximately 3.5 ppm, for example.
Herein, the signal that a nuclear spin that is present in fat tissue emits upon relaxation is designated as a “fat signal”. The signal that a nuclear spin that is located in an aqueous region emits upon relaxation is designated as an “aqueous signal” (“water signal”).
In many cases in the diagnosis of possible pathologies, it is problematic that the very bright fat signal outshines the aqueous signal that is of primary interest.
Therefore, possibilities have already been proposed to suppress the fat signal. For example, before the actual measurement a frequency-selective pulse is emitted at the precession frequency of the protons situated in the adipose tissue, such that their spins are saturated and no longer contribute to the signal in the subsequent image acquisition.
Due to eddy current effects and due to a limited bandwidth of the RF pulses of the B1 field (due to limited B1 peak, SAR etc.), a signal shadowing in the generated MRT images (for example in adipose tissue) can occur which can reduce the diagnostic value of the MRT images.
Eddy currents of an emitted or switched (activated) gradient widen the frequency distribution (spectrum) of the tissue within the region to be examined (FOV=field of view). This tissue is often a mixture of fat and water and, due to the chemical shift, as explained above there is a spacing between fat and water. This spacing becomes even wider due to the additional magnetic field (during the RF pulse) of the eddy currents. If the bandwidth of the RF pulses is limited, the entire spectrum cannot be refocused or, respectively, excited by the RF pulses. This problem of too little bandwidth in particular relates to the refocusing pulses, which normally are not spatially selective.
In more recent MRT systems, the available B1 peak of the RF pulses is reduced, such that this problem increasingly comes to the fore. This problem occurs less in a system with a B1 value of the RF pulses of 17 μT, but it occurs to an increased degree in a system with a B1 value of the RF pulses of 14 μT.